This invention relates generally to Computed Tomography (CT) imaging systems, and more particularly, to target angle heel effect compensation.
In at least some known imaging systems, an x-ray tube source projects an x-ray beam which passes through an object being imaged, such as a patient, and impinges upon an array of x-ray detector rows. The heel effect from an x-ray tube is well known. An x-ray tube source includes an anode side and a cathode side. The anode side is also known as the target, which is bombarded with electrons to generate x-ray beam radiation. X-rays from the x-ray tube are generated at a small depth inside the target (anode) of the x-ray tube. X-rays traveling toward an anode side of an object being scanned travel through more volume of the target than x-rays traveling toward a cathode side of the object. Therefore, x-rays traveling toward the anode side leave the target more attenuated than x-rays traveling toward the cathode side. This attenuation difference is called the heel effect.
With the advent of multislice CT imaging systems including a plurality of detector rows, the heel effect can produce image quality differences over the detector rows. For example, a 40 mm Volumetric Computed Tomography (VCT) detector with a nominal 7 degree target angle has an effective target angle of 5 degrees on the outer anode side row and 9 degrees on the outer cathode side row, resulting in a 20% effective signal difference between the anode side and the cathode side. This variance in radiation intensity due to the heel effect reduces image quality over the x-ray detector rows, and therefore reduces the image quality of the radiographs.